Manufacturing process for polymeric stents

ABSTRACT

Methods and systems of fabricating a polymeric stent are disclosed herein.

This application is a divisional of U.S. application Ser. No. 14/255,896file Apr. 17, 2014, which is a continuation of U.S. application Ser. No.14/252,689 filed Apr. 14, 2014, now U.S. Pat. No. 9,198,782, which is acontinuation of U.S. application Ser. No. 13/112,960 filed May 20, 2011,which is a continuation of U.S. application Ser. No. 11/443,947 filed onMay 30, 2006, now U.S. Pat. No. 7,971,333, all of which are incorporatedby reference herein.

BACKGROUND OF THE INVENTION

Field of the Invention

This invention relates to methods of manufacturing polymeric stents.

Description of the State of the Art

This invention relates to radially expandable endoprostheses, that areadapted to be implanted in a bodily lumen. An “endoprosthesis”corresponds to an artificial device that is placed inside the body. A“lumen” refers to a cavity of a tubular organ such as a blood vessel. Astent is an example of such an endoprosthesis. Stents are generallycylindrically shaped devices that function to hold open and sometimesexpand a segment of a blood vessel or other anatomical lumen such asurinary tracts and bile ducts. Stents are often used in the treatment ofatherosclerotic stenosis in blood vessels. “Stenosis” refers to anarrowing or constriction of a bodily passage or orifice. In suchtreatments, stents reinforce body vessels and prevent restenosisfollowing angioplasty in the vascular system. “Restenosis” refers to thereoccurrence of stenosis in a blood vessel or heart valve after it hasbeen treated (as by balloon angioplasty, stenting, or valvuloplasty)with apparent success.

Stents are typically composed of scaffolding that includes a pattern ornetwork of interconnecting structural elements or struts, formed fromwires, tubes, or sheets of material rolled into a cylindrical shape.This scaffolding gets its name because it physically holds open and, ifdesired, expands the wall of the passageway. Typically, stents arecapable of being compressed or crimped onto a catheter so that they canbe delivered to and deployed at a treatment site.

Delivery includes inserting the stent through small lumens using acatheter and transporting it to the treatment site. Deployment includesexpanding the stent to a larger diameter once it is at the desiredlocation. Mechanical intervention with stents has reduced the rate ofrestenosis as compared to balloon angioplasty. Yet, restenosis remains asignificant problem. When restenosis does occur in the stented segment,its treatment can be challenging, as clinical options are more limitedthan for those lesions that were treated solely with a balloon.

Stents are used not only for mechanical intervention but also asvehicles for providing biological therapy. Biological therapy usesmedicated stents to locally administer a therapeutic substance.Effective concentrations at the treated site require systemic drugadministration which often produces adverse or even toxic side effects.Local delivery is a preferred treatment method because it administerssmaller total medication levels than systemic methods, but concentratesthe drug at a specific site. Local delivery thus produces fewer sideeffects and achieves better results.

A medicated stent may be fabricated by coating the surface of either ametallic or polymeric scaffolding with a polymeric carrier that includesan active or bioactive agent or drug. Polymeric scaffolding may alsoserve as a carrier of an active agent or drug.

The stent must be able to satisfy a number of mechanical requirements.The stent must be capable of withstanding the structural loads, namelyradial compressive forces, imposed on the stent as it supports the wallsof a vessel. Therefore, a stent must possess adequate radial strength.Radial strength, which is the ability of a stent to resist radialcompressive forces, is due to strength and rigidity around acircumferential direction of the stent. Radial strength and rigidity,therefore, may also be described as, hoop or circumferential strengthand rigidity.

Once expanded, the stent must adequately maintain its size and shapethroughout its service life despite the various forces that may come tobear on it, including the cyclic loading induced by the beating heart.For example, a radially directed force may tend to cause a stent torecoil inward. In addition, the stent must possess sufficientflexibility to allow for crimping, expansion, and cyclic loading.

Some treatments with implantable medical devices require the presence ofthe device only for a limited period of time. Once treatment iscomplete, which may include structural tissue support and/or drugdelivery, it may be desirable for the stent to be removed or disappearfrom the treatment location. One way of having a device disappear may beby fabricating the device in whole or in part from materials that erodeor disintegrate through exposure to conditions within the body. Thus,erodible portions of the device can disappear or substantially disappearfrom the implant region after the treatment regimen is completed. Afterthe process of disintegration has been completed, no portion of thedevice, or an erodible portion of the device will remain. In someembodiments, very negligible traces or residue may be left behind.Stents fabricated from biodegradable, bioabsorbable, and/or bioerodiblematerials such as bioabsorbable polymers can be designed to completelyerode only after the clinical need for them has ended.

However, there are potential shortcomings in the use of polymers as amaterial for implantable medical devices, such as stents. There is aneed for a manufacturing process for a stent that addresses suchshortcomings so that a polymeric stent can meet the clinical andmechanical requirements of a stent.

SUMMARY OF THE INVENTION

Certain embodiments of the invention include a method of fabricating astent delivery device comprising forming a polymeric tube usingextrusion, the tube being drawn during extrusion so that a diameter ofthe formed tube is less than a target diameter; radially deforming theformed tube so that the deformed tube comprises the target diameter;forming a stent from the deformed tube, wherein forming the stentincludes laser machining a stent pattern in the deformed tube with anultra-short pulse laser; and crimping the stent on a support element,wherein a temperature of the stent during crimping is above an ambienttemperature.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 depicts a stent.

FIG. 2 depicts an axial cross-section of a polymer tube positionedwithin an annular member or mold.

FIG. 3 depicts a deformed polymer tube in a mold.

FIG. 4 depicts a schematic plot of the crystal nucleation rate and thecrystal growth rate, and the overall rate of crystallization.

FIG. 5 depicts a mathematical representation of a Gaussian laser beamprofile.

FIG. 6 depicts a collimated two-dimensional representation of a laserbeam.

FIG. 7 depicts an overhead view of the surface of a substrate.

FIG. 8 illustrates a kerf machined by a laser.

FIG. 9 depicts an exemplary stent pattern.

FIG. 10A depicts an exemplary bending element of a stent pattern.

FIG. 10B depicts the strain distribution along a line offset from theneutral axis of the bending element depicted in FIG. 10B.

FIG. 11 is a schematic plot of the specific volume of an amorphouspolymer vs. temperature.

DETAILED DESCRIPTION OF THE INVENTION

Various embodiments of methods of manufacturing a polymeric stent and adelivery system for a stent are disclosed. A stent can be made of abiostable, biodegradable, or a combination of a biostable andbiodegradable polymer. The terms degrade, absorb, and erode, as well asdegraded, eroded, and absorbed, are used interchangeably and refer tomaterials that are capable of being completely eroded, or absorbed whenexposed to bodily conditions.

A stent may include a pattern or network of interconnecting structuralelements or struts. FIG. 1 depicts an example of a three-dimensionalview of a stent 10. The stent may have a pattern that includes a numberof interconnecting elements or struts 15. The embodiments disclosedherein are not limited to stents or to the stent pattern illustrated inFIG. 1.

The embodiments described herein are easily applicable to other patternsand other implantable medical devices, including, but not limited toself-expandable stents, balloon-expandable stents, stent-grafts, andgrafts. The structural pattern of the device can be of virtually anydesign. The variations in the structure of patterns are virtuallyunlimited. As shown in FIG. 1 the geometry or shape of stents varythroughout its structure. A pattern may include portions of struts thatare straight or relatively straight, an example being a section 20.Patterns may also include intersections of struts with curved or bentportions or elements as in sections 25 and 30. In addition, patterns mayinclude struts that include curved or bent portions or elements as in asection 35.

Additionally, a surface of a medical device, such as a stent, may alsobe characterized by the relative location of the surface with respect toa bodily lumen. The stent includes abluminal surfaces or outer portions,luminal surfaces or inner portions, and surfaces between the abluminaland luminal surfaces. For example, struts 15 of stent 10 include luminalfaces or surfaces 40, abluminal faces or surfaces 45, and side-wallfaces or surfaces 50.

A stent such as stent 10 may be fabricated from a polymeric tube or asheet by rolling and bonding the sheet to form the tube. A tube or sheetcan be formed by extrusion or injection molding. A stent pattern, suchas the one pictured in FIG. 1, can be formed in a tube or sheet with atechnique such as laser cutting or chemical etching. The stent can thenbe crimped on to a balloon or catheter for delivery into a bodily lumen.

A stent made from a biodegradable polymer is intended to remain in thebody for a duration of time until its intended function of, for example,maintaining vascular patency and/or drug delivery is accomplished. Afterthe process of degradation, erosion, absorption, and/or resorption hasbeen completed, no portion of the biodegradable stent, or abiodegradable portion of the stent will remain. In some embodiments,very negligible traces or residue may be left behind.

The duration of a treatment period depends on the bodily disorder thatis being treated. In treatments of coronary heart disease involving useof stents in diseased vessels, the duration can be in a range from abouta month to a few years. However, the duration is typically up to aboutsix months, twelve months, eighteen months, or two years. In somesituations, the treatment period can extend beyond two years. Stentscomposed of biodegradable polymers are particularly attractive sincethey can be designed to remain in the body during the above-mentionedtime frames.

In addition, to biodegradability, there are several properties that aregenerally desirable for a stent to have that greatly facilitate thedelivery, deployment, and treatment of a diseased vessel. As indicatedabove, a stent has certain mechanical requirements. A stent must havesufficient radial strength to withstand structural loads, namely radialcompressive forces, imposed on the stent as it supports the walls of avessel. In addition, the stent must possess sufficient flexibility toallow for crimping, expansion, and cyclic loading. Also, a sufficientlylow profile, that includes diameter and size of struts, is important. Asthe profile of a stent decreases, the easier is its delivery, and thesmaller the disruption of blood flow.

Polymers tend to have a number of shortcomings for use as materials forstents. One such shortcoming is that many biodegradable polymers have arelatively low modulus, and thus relatively low radial strength.Compared to metals, the strength to weight ratio of polymers is smallerthan that of metals. A polymeric stent with inadequate radial strengthcan result in mechanical failure or recoil inward after implantationinto a vessel. To compensate for the relatively low modulus, a polymericstent requires significantly thicker struts than a metallic stent, whichresults in an undesirably large profile.

Another shortcoming of polymers is that many polymers, such asbiodegradable polymers, tend to be brittle under biological conditionsor conditions within a human body. Specifically, such polymers can havea Tg above human body temperature which is approximately 37° C. Thesepolymer systems exhibit a brittle fracture mechanism in which there islittle or no plastic deformation prior to failure. As a result, a stentfabricated from such polymers can have insufficient toughness for therange of use of a stent. In particular, it is important for a stent tobe resistant to fracture throughout the range of use of a stent, i.e.,crimping, delivery, deployment, and during a desired treatment period.

Other potential problems with polymeric stents include creep, stressrelaxation, and physical aging. Creep refers to the gradual deformationthat occurs in a polymeric construct subjected to an applied load. Creepoccurs even when the applied load is constant.

It is believed that the delayed response of polymer chains to stressduring deformation causes creep behavior. As a polymer is deformed,polymeric chains in an initial state rearrange to adopt a newequilibrium configuration. Rearrangement of chains takes place slowlywith the chains retracting by folding back to their initial state. Forexample, an expanded stent can retract radially inward, reducing theeffectiveness of a stent in maintaining desired vascular patency. Therate at which polymers creep depends not only on the load, but also ontemperature. In general, a loaded construct creeps faster at highertemperatures.

Stress relaxation is also a consequence of delayed molecular motions asin creep. Contrary to creep, however, which is experienced when the loadis constant, stress relaxation occurs when deformation (or strain) isconstant and is manifested by a reduction in the force (stress) requiredto maintain a constant deformation. The mechanism responsible for bothstress relaxation and creep will result in recoil, as the stent mayexperience deformation in response to an applied load (from the vessel).Unlike stress relaxation, the deformation imposed on the stent maychange with time. Unlike creep, the applied load will also change withtime.

Physical aging, as used herein, refers to densification in the amorphousregions of a semi-crystalline polymer. Densification is the increase indensity of a material or region of a material. Densification resultsfrom residual and applied stresses. As used herein, a “residual stress”includes, without limitation, the stress in a bulk polymer that is in anon-equilibrium thermodynamic state. Physical aging of semi-crystallinepolymers that have glass transition temperatures (Tg) above their normalstorage temperature, which, for the purposes of this invention is roomtemperature, i.e., from about 15° C. to about 35° C., occurs primarilythrough the phenomenon known as densification.

Densification occurs when a semi-crystalline polymer is cooled at anon-equilibrium rate from a temperature above its Tg to a temperaturebelow its Tg. Non-equilibrium cooling is normally what will occur inmost industrial settings since equilibrium cooling is very slow andwould be considered economically impractical. The non-equilibriumcooling rate results in the polymer chains of the amorphous domainsbeing trapped at non-optimal separation distances in the glassy statethat forms when the temperature goes below Tg. The chains then attemptto achieve optimal separation by coordinated localized chain motion. Thereordering of polymer chains tends to increase the modulus of thepolymer resulting in a brittle or more brittle polymer. Thus,densification of a polymer initially selected for its toughness andelasticity could cause failure of a coating or polymeric scaffoldingwhen the polymer ages or densifies and becomes brittle.

Certain embodiments of the present invention relate to methods offabricating a polymeric stents that have one or more above-mentionedrequired or desirable properties including biodegradability, adequatetoughness in the conditions of use of a stent, and a strength to weightratio close to metals resulting in an acceptable stent profile. Theembodiments can also reduce or eliminate above mentioned shortcomings ofpolymeric materials such as creep, stress relaxation, and physicalaging.

In certain embodiments, a method of forming a stent may include forminga tube and laser machining a tube to form a stent pattern in a mannerthat the formed stent has the above-mentioned required or desirableproperties. Certain embodiments of a method of fabricating a stent mayinclude:

(1) forming a polymeric tube using extrusion,

(2) radially deforming the formed tube, and

(3) forming a stent from the deformed tube by laser machining a stentpattern in the deformed tube with an ultra-short pulse laser.

In some embodiments, the tubing can be extruded to dimensions that allowradial deformation of the tube to a desired size for fabrication of astent. Radial deformation tends to increase the radial strength of thetubing, and the subsequently fabricated stent. Additionally, the tubecan be extruded so that the resultant tube has relatively lowcrystallinity which reduces brittle behavior and increases toughness.

Other embodiments of a method of fabricating a polymeric stent caninclude one or more of the following of the additional features:

(1) the method can include the use of an ultrafast laser reduces thedeleterious affects on mechanical properties of the polymer during lasermachining,

(2) the method can include forming a stent pattern in which the strainexperienced by the stent is spread or distributed over a relativelylarge area rather than being highly concentrated when the stent isdeformed, and

(4) embodiments can include fabricating the stent from a semicrystallinepolymer which can reduce or eliminate creep, stress relaxation, andphysical aging.

Further embodiments can include fabricating a stent delivery device bycrimping the stent on a support element such that the temperature of thestent during crimping is above an ambient temperature. Heating a stentduring crimping can reduce or eliminate radially outward recoiling of acrimped stent which can result in an unacceptable profile for delivery.

Polymer tubes may be formed using various types of forming methods,including, but not limited to extrusion or injection molding.Alternatively, a polymer tube may be formed from sheets that are rolledand bonded into a tube. Representative examples of extruders include,but are not limited to, single screw extruders, intermeshing co-rotatingand counter-rotating twin-screw extruders and other multiple screwmasticating extruders.

In extrusion, a polymer melt is conveyed through an extruder and forcedthrough a die in the shape of as an annular film in the shape of a tube.The annular film can be cooled below the melting point, Tm, of thepolymer to form an extruded polymeric tube.

In one embodiment, the annular film in the shape of a tube can beaxially drawn or stretched. As the tube is drawn, its diameterdecreases. Therefore, the tube can be drawn to have a selected diameter.In some embodiments, the annular film may be drawn by a puller. Thepuller may include a conveyor assembly that supports and sizes theannular film.

The annular film may be cooled during expansion and/or after drawing.For example, the annular film may be conveyed through a water bath at aselected temperature. Alternatively, the annular film may be cooled byair at a selected temperature. The annular film may be cooled at or nearan ambient temperature, e.g. 25° C. Alternatively, the annular film maybe cooled at a temperature below ambient temperature.

As mentioned above, polymers tend of have a relatively low strength toweight ratio as compared to metals. Fabrication of a polymeric stentwith strength comparable to a conventional metallic stent can requireundesirably large struts.

Thus, in certain embodiments, a method of fabricating a stent mayinclude modifying the mechanical properties of polymer construct, suchas a tube, to have desirable mechanical properties. The polymer can bemodified to increase the strength, modulus, and/or toughness of thepolymer tube. In some embodiments, the polymer can be modified to allowfabrication of a polymeric stent with strength comparable to aconventional metallic stent and with strut dimensions that are close tothe dimensions of conventional metallic stents.

It is well known by those skilled in the art that the mechanicalproperties of a polymer can be modified by applying stress to a polymer.The strength and modulus of a polymer can be increased along thedirection of the applied stress. The application of stress can inducemolecular orientation along the direction of stress which can increasethe strength and modulus along the direction. Molecular orientationrefers to the relative orientation of polymer chains along alongitudinal or covalent axis of the polymer chains.

Molecular orientation can be induced in polymers that are completelyamorphous, partially or semi-crystalline, or almost completelycrystalline. A partially or semi-crystalline polymer includescrystalline regions separated by amorphous regions. The crystallineregions do not necessarily have the same or similar orientation ofpolymer chains. However, a high degree of orientation of crystallitesmay be induced by applying stress to a semi-crystalline polymer. Thestress may also induce orientation in the amorphous regions.

Due to the magnitude and directions of stresses imposed on a stentduring use, it is important for the mechanical stability of a device tohave an adequate magnitude of strength both in axial and circumferentialdirections. Therefore, an adequate balance of axial and circumferentialstrength is also important for mechanical stability. The relative amountof axial and circumferential orientation may depend on a number offactors such as the stent pattern, initial diameter of the tube, finaldiameter of the stent, and crimped diameter of the stent. Polymer tubesformed by extrusion methods tend to possess a significant degree ofaxial polymer chain alignment. However, such conventionally extrudedtubes tend to possess no or substantially no polymer chain alignment inthe circumferential direction.

Some embodiments of a method of fabricating a stent may include radiallydeforming a polymeric tube about a cylindrical axis of the tube. Thetube can be radially deformed to increase the strength and modulus inthe circumferential direction. The increase in strength and modulus canbe due to the induced molecular orientation in the circumferentialdirection.

Additionally, the method may further include axially deforming the tubealong the cylindrical axis of the tube. In one embodiment, the tube maybe axially deformed by applying a tensile force to the tube along thecylindrical axis. Axial deformation of the polymer tube may induce axialmolecular orientation, and hence, increase the axial strength andmodulus or rigidity. Various embodiments may include radially deformingthe tube prior to, subsequent to, and/or contemporaneously with axialdeformation the tube.

The degree of polymer chain alignment induced with applied stress maydepend upon the temperature of the polymer. Above Tg, polymer chainalignment may be readily induced with applied stress since polymerchains have greater mobility than below Tg. Consequently, the amount ofdeformation depends on the temperature of a polymeric material.Therefore, it is advantageous to radially deform the tube at atemperature above Tg.

Additionally, the polymeric tube may be heat set to allow polymericchains to rearrange upon deformation. “Heat setting,” as used herein,refers to maintaining a polymer at an elevated temperature to allowpolymer chains in the heated polymer to move toward a state ofthermodynamic equilibrium. In a deformed polymeric tube, polymericchains are allowed to equilibrate towards the induced highly orientedstructure at the elevated temperature. Since polymer chain alignment isa time and temperature dependent process, a highly oriented structurethat is thermodynamically stable at a given temperature may not beformed instantaneously. Thus, the polymeric tube may be maintained in adeformed state at an elevated temperature for a period of time.

A tube can be radially deformed using blow molding. FIGS. 2 and 3illustrate an embodiment of deforming a polymeric tube in manufacturinga stent. FIG. 2 depicts an axial cross-section of a polymeric tube 150with an outside diameter 155 positioned within an annular member or mold160. Mold 160 may act to limit the radial deformation of polymeric tube150 to a diameter 165, the inside diameter of mold 160. Polymer tube 150may be closed at a distal end 170. Distal end 170 may be open insubsequent manufacturing steps. A fluid may be conveyed, as indicated byan arrow 175, into an open proximal end 180 of polymeric tube 150. Atensile force 195 is applied at proximal end 180 and a distal end 170.

Polymeric tube 150 may be heated by heating the gas to a temperatureabove ambient temperature prior to conveying the gas into polymeric tube150. Alternatively, the polymeric tube may be heated by heating theexterior of mold 160. The tube may also be heated by the mold. Theincrease in pressure inside of polymer tube 150 facilitated by anincrease in temperature of the polymeric tube causes radial deformationof polymer tube 150, as indicated by an arrow 185. FIG. 3 depictspolymeric tube 150 in a deformed state with an outside diameter 190within annular member 160.

Additionally, as indicated above, the pressure inside the tube and thetemperature of the tube may be maintained at the elevated temperaturefor a period of time to allow the polymeric tube to be heat set. Theperiod of time may be between about one minute and about two hours, ormore narrowly, between about two minutes and about ten minutes.

Furthermore, the tube may be expanded to a target diameter. In oneembodiment, the target diameter may be the diameter at which a stentpattern is formed by laser machining the tube. The target diameter canalso correspond to the diameter of a stent prior to crimping. The degreeof radial deformation may be quantified by a blow-up ratio or radialdraw ratio:

$\frac{{Outside}\mspace{14mu} {Diameter}\mspace{14mu} {of}\mspace{14mu} {Deformed}\mspace{14mu} {Tube}}{{Original}\mspace{14mu} {Inside}\mspace{14mu} {Diameter}\mspace{14mu} {of}\mspace{14mu} {Tube}}$

In some embodiments, the radial draw ratio of a polymeric tube for usein fabricating a stent may be between about 1 and 20, or more narrowlybetween about 2 and 6. Similarly, the degree of axial deformation may bequantified by an axial draw ratio:

$\frac{{Length}\mspace{14mu} {of}\mspace{14mu} {Deformed}\mspace{14mu} {Tube}}{{Original}\mspace{14mu} {Length}\mspace{14mu} {of}\mspace{14mu} {Tube}}$

In some embodiments, a stent can be fabricated from the radiallyexpanded tube by laser machining. The stent may then be crimped on to adelivery device such as a balloon. Therefore, the extruded tube musthave a diameter that is less than a target diameter, the target diametercorresponding to a diameter at which a stent pattern is formed by lasermachining or a diameter of a stent prior to crimping. Thus, the extrudedtube can then be radially expanded to the target diameter.

The degree of induced strength due to radial expansion depends on theamount of radial expansion as quantified by the blow-up ratio. Thus, thedegree of radial expansion is determined by the extruded diameter andthe target diameter. In some embodiments, the extruded diameter can beused to determine the degree of radial expansion. The smaller theextruded diameter, the larger is the degree of radial expansion. In someembodiments, the extruded diameter can be controlled by the drawing ofthe polymer film as it exits the die. For a die having a diametersimilar to the target diameter, the axial draw down ratio can be between1 and 3. The axial drawdown ratio is defined as:

$\frac{{Drawn}\mspace{14mu} {Length}\mspace{14mu} {of}\mspace{14mu} {Tube}}{{Original}\mspace{14mu} {Length}\mspace{14mu} {of}\mspace{14mu} {Tube}}$

In further embodiments, it may be desirable to control the degree ofcrystallinity in a polymer tube during extrusion, radial deformation,and/or axial deformation to reduce or eliminate physical aging, creep,and stress relaxation in a stent. Additionally, it is also desirable tocontrol crystallinity to increase the toughness of a stent.

As discussed above, physical aging, creep, and stress relaxation are dueat least in part to rearrangement of polymer chains in amorphous regionsof a polymer. Thus, as crystallinity increases in a polymer, physicalaging, creep, and stress relaxation tend to reduced. Therefore, it isadvantageous to use crystalline or semi-crystalline polymers for a stentto reduce or eliminate physical aging, creep, and stress relaxation.

However, crystalline and semi-crystalline polymers can be relativelybrittle at biological conditions, i.e., the temperature of the humanbody. In particular, such polymers can have a Tg below the bodytemperature. These polymers can have a low fracture toughness and arethus susceptible to mechanical failure during use, for example, duringcrimping, deployment, and treatment. It is important for a stent to havea high fracture toughness throughout the range of stress experiencedduring use. For a semi-crystalline polymer, the higher thecrystallinity, the more likely the polymer will be brittle underbiological conditions.

Semi-crystalline polymers can contain both amorphous and crystallinedomains at temperatures below their melting point. Amorphous regions arethose in which polymer chains are in relatively disorderedconfigurations. Crystalline domains are those in which polymer chainsare in ordered configurations with segments of polymer chainsessentially parallel to one another.

In certain embodiments, the crystallinity of an extruded tube can becontrolled by controlling the temperature of cooling the annular filmexiting the die. In general, crystallization tends to occur in a polymerat temperatures between Tg and Tm of the polymer. The rate ofcrystallization in this range varies with temperature. FIG. 4 depicts aschematic plot of the crystal nucleation rate (R_(N)), the crystalgrowth rate (R_(CG)), and the overall rate of crystallization (R_(CO)).The crystal nucleation rate is the growth rate of new crystals and thecrystal growth rate is the rate of growth of formed crystals. Theoverall rate of crystallization is the sum of curves R_(N) and R_(CG).

In certain embodiments, the temperature of the annular tube exiting theextruder during cooling can be at a temperature at which the overallcrystallization rate is relatively low. In some embodiments, thetemperature can be in a range in which the crystal nucleation rate islarger than the crystal growth rate. In one embodiment, the temperaturecan be in a range in which the crystal nucleation rate is substantiallylarger than the crystal growth rate. For example, the temperature can bewhere the ratio of the crystal nucleation rate to crystal growth rate is2, 5, 10, 50, 100, or greater than 100. In another embodiment, thetemperature range may be in range, ΔT shown in FIG. 4, between about Tgto about 0.5(Tm−Tg)+Tg.

The resulting polymeric tube may then have relatively low crystallinity.Additionally, since the nucleation rate is higher than the crystalgrowth rate, the resulting polymer can have a relatively large number ofcrystalline domains that are relatively small. As the size of thecrystalline domains decreases along with an increase in the number ofdomains, the polymer may become less brittle and, thus, which increasesthe fracture toughness.

Although the crystallinity of the resulting polymer can be relativelylow, the presence of the relatively large number of relatively smallcrystalline domains can reduce or eliminate physical aging, creep, andstress relaxation. It is believed that crystallites inhibit movement ofpolymer chains in the amorphous regions. As indicated above, suchmovement contributes to physical aging, creep, and stress relaxation.

Additionally, crystallinity in a polymeric tube can also be controlledduring radial deformation to increase fracture toughness and reducephysical aging, creep, and stress relaxation. As indicated above, it isdesirable to radial deform at a temperature above a Tg of the polymer tofacilitate deformation and to heat set the polymer above the Tg. As aresult, crystallization tends to occur in the polymer during deformationand heat setting. In addition, crystallinity in addition to orientationmay be induced in the polymer during deformation. This process isreferred to as strain-induced crystallization.

Thus, embodiments of the method can include radially deforming and/orheat setting in temperature ranges described above for extrusion.Deforming and heat setting a tube in such a temperature range can resultin a radially deformed tube with a higher fracture toughness and reducedphysical aging, creep, and stress relaxation for the same reasons asexplained above for extrusion.

In other embodiments, an extruded tube can be formed that is amorphousor substantially amorphous. An amorphous polymeric tube can be formed byquenching the annular tube exiting the die so that its temperature isreduced to a temperature from above Tm to below Tg so that no orsubstantially no crystallization occurs in the polymer during cooling.Thus, an amorphous glassy polymer can be formed. The tube can then bedeformed and heat set, as described above, at a temperature that resultsin high fracture toughness and with reduced physical aging, creep, andstress relaxation.

Additionally, certain embodiments of a method of fabricating a stent caninclude laser machining a stent pattern in an extruded and radiallydeformed tube. In some embodiments, a stent pattern may be formed withan ultrashort-pulse lasers. “Ultrashort-pulse lasers” refer to lasershaving pulses with durations shorter than about a picosecond (=10⁻¹²).Ultrashort-pulse lasers can include both picosecond and femtosecond(=10⁻¹⁵) lasers. The ultrashort-pulse laser is clearly distinguishablefrom conventional continuous wave and long-pulse lasers (nanosecond(10⁻⁹) laser) which have significantly longer pulses. Certainembodiments of the present method may employ femtosecond lasers havingpulses shorter than about 10⁻¹³ second.

The ultrashort-pulse lasers are known to artisans. For example, they arethoroughly disclosed by M. D. Perry et al. in Ultrashort-Pulse LaserMachining, Section K-ICALEO 1998, pp. 1-20. Representative examples offemtosecond lasers include, but are not limited to a Ti:sapphire laser(735 nm-1035 nm) and an excimer-dye laser (220 nm-300 nm, 380 nm-760nm).

A significant advantage of ultra-short lasers as compared to longerpulse lasers is that ultra-short lasers tend to result in asignificantly smaller heat affected zone (HAZ). A HAZ refers to aportion of a target substrate that is not removed, but is still exposedto energy from the laser beam, either directly or indirectly. Directexposure may be due to exposure to the substrate from a section of thebeam with an intensity that is not great enough to remove substratematerial through either a thermal or nonthermal mechanism. For example,the portions of a beam near its edges may not have an intensitysufficiently high to remove substrate material. Most laser beams have aspatially uneven or nonuniform beam intensity profile, for example, aGaussian beam profile. A substrate can also be exposed to energyindirectly due to heat conduction.

A heat affected zone in a target substrate is undesirable for a numberof reasons. In polymers, heat can cause thermal distortion and roughnessat the machined surface. The heat can also alter properties of a polymersuch as mechanical strength and degradation rate. The heat can alsocause chemical degradation that can affect the mechanical properties anddegradation rate.

Additionally, heat can modify molecular structure of a polymer, such asdegree of crystallinity and polymer chain alignment. Mechanicalproperties are highly dependent on molecular structure. As mentionedabove, a high degree of crystallinity and/or polymer chain alignment isassociated with a stiff, high modulus material. Heating a polymer aboveits melting point can result in an undesirable increase or decrease incrystallinity once the polymer resolidifies. Melting a polymer may alsoresult in a loss of polymer chain alignment induced by radial expansion,which can adversely affect mechanical properties.

In addition, since heat from the laser modifies the properties of thesubstrate locally, the mechanical properties may become spatiallynonuniform. Such nonuniformity may lead to mechanical instabilities suchas cracking.

Longer-pulse lasers remove material from a surface principally through athermal mechanism. The laser energy that is absorbed results in atemperature increase at and near the absorption site. As the temperatureincreases to the melting or boiling point, material is removed byconventional melting or vaporization. Depending on the pulse duration ofthe laser, the temperature rise in the irradiated zone may be very fast,resulting in thermal ablation and shock. An advantage ofultrashort-pulse lasers over longer-pulse lasers is that theultrashort-pulse deposits its energy so fast that is does not interactwith the plume of vaporized material, which would distort and bend theincoming beam and produce a rough-edged cut.

Unlike long-pulse lasers, ultrashort-pulse lasers allow material removalby a nonthermal mechanism. Extremely precise and rapid machining can beachieved with essentially no thermal ablation and shock. The nonthermalmechanism involves optical breakdown in the target material whichresults in material removal. As discussed below, optical breakdown mayalso occur with a gas, in particular with a process gas in contact withthe beam at the surface of the target substrate. Optical breakdown tendsto occur at a certain threshold intensity of laser radiation that ismaterial dependent. Specifically each material has its own laser-inducedoptical breakdown threshold which characterizes the intensity requiredto ablate the material at a particular pulse width.

During optical breakdown of material, a very high free electron density,i.e., plasma, is produced. The plasma can be produced through mechanismssuch as multiphoton absorption and avalanche ionization.

In optical breakdown, a critical density plasma is created in a timescale much shorter than electron kinetic energy is transferred to thesubstrate lattice. The resulting plasma is far from thermal equilibrium.The target material is converted from its initial solid-state directlyinto a fully ionized plasma on a time scale too short for thermalequilibrium to be established with a target material lattice. Therefore,there is negligible heat conduction beyond the region removed. As aresult, there is negligible thermal stress or shock to the materialbeyond approximately 1 micron from the laser machined surface.

In laser machining with longer-pulse and ultra-fast pulse lasers,material removal tends to occur in an area or region of directinteraction of a laser beam with the target material or substrate. Lasermachining typically involves focusing a laser beam onto an area orregion of the substrate. The area of direct interaction corresponds to afocus diameter (D_(f)) on the target material that can be calculatedfrom:

D _(f)=1.27*f*λ/D

where f is the focal length of a focusing optic, λ is the wave length ofthe laser, and D is the beam diameter on the optic.

FIG. 5 depicts an axial cross-section of a laser beam 501 traveling inthe “z” direction as indicated by an arrow 502. A mathematicalrepresentation 504 in the form of a Gaussian beam profile is shownsuperimposed on the beam. The profile has a maximum intensity (I_(max))at the beam center (x=0) and then decreases with distance on either sideof the maximum. The sections of the beam close to the edge may notremove material. However, such sections may still deposit energy intothe material that can have undesirable thermal affects. Additionally, aportion of the substrate may also be heated through conduction.

FIGS. 6-8 are schematic illustrations of laser machining a substrate.FIG. 6 depicts a collimated two-dimensional representation of a laserbeam 610 passing through a focusing lens 612 with a focal point 614. Afocused laser beam 616 decreases in diameter with distance from lens612. Beam 616 impinges on a substrate 618. The region of directinteraction of the laser beam has a diameter D_(f).

FIG. 7 depicts an overhead view of the surface of substrate 618 showingregion 720 with a diameter D_(D) which corresponds to an area in whichmaterial is removed from the substrate. Region 720 is smaller than thearea of direct interaction of the laser which has a diameter D_(f). FIG.8 illustrates that translation of the laser beam or substrate allows thelaser beam to cut a trench or kerf 824 with at least a width D_(D).Energy is deposited in region 828, but does not remove material. Region828 corresponds to a portion of a heat affected zone. The heat affectedzone can extend beyond region 828 since substrate material can also beheated by heat conduction.

Assemblies for laser machining stents have been described in numerouspatents including U.S. Pat. Nos. 6,521,865 and 6,131,266. Suchassemblies can readily be modified for use with an ultra-fast pulselaser. Generally, a laser beam source may be positioned to direct thebeam from the laser beam source to remove material from a substrate heldby a fixture. In some embodiments, a polymeric tube can be disposed in arotatable collet fixture of a machine-controlled apparatus thatpositions the tubing relative to a laser. According to machine-encodedinstructions, the tube can be rotated and moved axially relative to thelaser which can also be machine-controlled. The laser selectivelyremoves the material from the tubing resulting in a pattern cut into thetube. The tube is therefore cut into the discrete pattern of a finishedstent.

In certain embodiments, the process of cutting a pattern for the stentinto the tubing can be automated except for loading and unloading thelength of tubing. The automation may be done, for example, using aCNC-opposing collet fixture for axial rotation of the length of tubing.A collet fixture may act in conjunction with a CNC X/Y table to move thelength of tubing axially relative to a machine-controlled laser asdescribed. The program for control of the apparatus is dependent on theparticular configuration used and the pattern formed. CNC equipmentmanufactured and sold by Anorad Corporation in Hauppauge, N.Y. may beused for positioning the tube.

It is desirable for a stent to have a stent pattern which allows thestrain experienced by the stent to be spread or distributed over arelatively large area rather than being highly concentrated when thestent is deformed. During crimping and expansion, curved regionsconfigured to bend can experience a strain that is localized. Suchregions of the stent can be the most susceptible to fracture.

FIG. 9 depicts an exemplary stent pattern 900 for use with embodimentsof a polymeric tube or a sheet. In an embodiment, stent pattern 900 canbe cut from a polymeric tube using embodiments of the laser machiningprocess described herein. The polymeric tube can be a radially and/oraxially expanded tube, as described above.

Stent pattern 900 is shown in a flattened condition so that the patterncan be clearly viewed. When the flattened portion of stent pattern 900is in a cylindrical form, it forms a radially expandable stent.

As depicted in FIG. 9, stent pattern 900 includes a plurality ofcylindrical rings 902 with each ring including a plurality of diamondshaped cells 904. Embodiments of stent pattern 900 may have any numberof rings 902 depending a desired length of a stent. For reference, lineA-A represents the longitudinal axis of a stent using the patterndepicted in FIG. 9. Diamond shaped cells 904 are made up of bar arms 906and 908 that form a curved element and bar arms 910 and 912 that form anopposing curved element.

Pattern 900 further includes linking arms 916 that connect adjacentcylindrical rings. Linking arms 916 are parallel to line A-A and connectadjacent rings between intersection 918 of cylindrically adjacentdiamond-shaped elements 904 of one ring and and intersection 918 ofcylindrically adjacent diamond shaped elements 904 of an adjacent ring.As shown, linking elements connect every other intersection along thecircumference.

Generally, a stent is allowed to be crimped or expanded through flexingof curved or bending elements. A stent with a pattern 900 can be crimpedor compressed principally through flexing of curved elements between

(1) bar arms within diamond-shaped cells symmetric about thelongitudinal axis bar arms, for example, bar arms 906 and 908 with anangle θ₁ (bending or curved element 1)

(2) bar arms of cylindrically adjacent diamond-shaped cells, forexample, bar arms 906 and 908 with an angle θ₂ (bending or curvedelement 2)

(3) a bar arm of a diamond-shaped cell and a linking bar arm 916, forexample, bar arms 908 and linking bar arm 916 with an angle θ₃ (bendingor curved element 3)

When a stent having a stent pattern 900 is crimped, bending elements 1,2, and 3 flex inward and angles θ₁, θ₂, and θ₃ decrease, allowing thestent to be radially compressed. Similarly, a stent is allowed to expandwhen angles θ₁, θ₂, and θ₃, increase. With respect to bending elements 1and 2, each of the bar arms of the respective bending elements tend tobend toward or away from each other. However, in bending element 3, thebar arm of the diamond-shaped element tends to bend toward the linkingbar arm, the linking bar arm which tends to remain relatively parallelto the longitudinal axis.

The region at and adjacent to the apices of the bending elements: apex920, 922, and 924, experience a high degree of stress and strain duringcrimping and expansion. The spatial distribution of the strain dependson the geometry of the bending portion. The strain can vary from beinghighly localized to being widely distributed in the apex regions.

FIG. 10A depicts an exemplary bending element 1000 and FIG. 10B depictsthe strain distribution along a line 1001 at the inner surface ofbending element 1000 that follows its curvature. Curve 1003 correspondsto a strain distribution that is highly localized while curve 1004represents a more disperse strain distribution. In general, a moredisperse strain distribution is more desirable since the more localizedthe strain, the more likely is mechanical failure at the region ofstrain concentration. As shown by curve 103, the maximum strainexperienced by bending arm 1000 increases the more concentrated thestrain distribution, resulting in a higher likelihood of fracture.

In general, the radius of curvature of a bending arm is directly relatedto the distribution of stress in a bending arm that is in either acrimped or compressed state. The larger the radius of curvature of anunstressed bending arm, the more widely dispersed or distributed is thestrain.

Therefore, the stress concentration in bending elements 1, 2, and 3 canbe controlled at least by the radius of curvature, R₁, R₂, and R₃,respectively, which are depicted in FIG. 10. In general, the radiishould be as large as possible while still allowing a stent to reach adesired crimped diameter.

An exemplary stent with a stent pattern 900 can have six diamond-shapedcells with a circumference of 0.264 in. In general, a stent with stentpattern 900 can have a ratio of its outer circumference to the number ofcells around the circumferences of 0.264 in/6 cells or 0.044 in/cell. Insome embodiments, a stent having a disperse strain distribution incurved elements can have R₁=0.0056 in, R₂=0.0034, and/or R₃=0.005. Inother embodiments, such a stent can have an R₁ between 0.004 in and0.006 in, R₂ between 0.0025 in and 0.004 in, and/or R₃ between 0.004 inand 0.006 in.

Embodiments of the stents above tend to have relatively dispersedistribution of strain in the apex regions of bending elements 1, 2, and3. For example, the magnitude of stress tends to be below the ultimatestrength of stents fabricated according to the embodiments describedherein.

As discussed above, prior to delivery into the body a stent iscompressed or crimped onto a catheter so that it can be inserted intosmall vessels. Once the stent is delivered to the treatment site, it canbe expanded or deployed at a treatment site. Generally, stent crimpingis the act of affixing the stent to the delivery catheter or deliveryballoon so that it remains affixed to the catheter or balloon until thephysician desires to deliver the stent at a treatment site. There arenumerous crimpers available for crimping stents including, but notlimited to, the roll crimper, collet crimper, and wedge crimper.

Conventionally, crimping is performed at ambient conditions. Ashortcoming of polymeric stents as compared with metallic stents is thatthey can be relatively brittle at ambient conditions. Such polymers havea Tg that is above ambient conditions. As discussed above, there areregions of a stent subjected to relatively high stress and strain duringcrimping. Thus, such regions are susceptible to mechanical failure insuch high strain regions due to the brittle nature of the polymer atambient conditions. Another shortcoming of polymeric stents with respectto crimping is that polymeric stents can have a tendency to recoiloutward from a crimped diameter. The recoiling can result in poor stentretention.

Certain embodiments of a crimping method for polymeric stents aredescribed herein which reduce or eliminate recoil and reduce oreliminate mechanical damage due to the crimping process. In someembodiments, a stent may be crimped at a temperature above ambientconditions that reduces or eliminates recoil and mechanical failure.

As discussed above, below Tg a polymer tends to be brittle and thus issusceptible to fracture at a relatively low elongation when subjected tostress. Generally, the glass transition temperature, Tg, is thetemperature at which the amorphous domains of a polymer change from abrittle vitreous state to a solid deformable or ductile state atatmospheric pressure. When an amorphous or semicrystalline polymer isexposed to an increasing temperature, the coefficient of expansion andthe heat capacity of the polymer both increase as the temperature israised, indicating increased molecular motion. As the temperature israised, the actual molecular volume in the sample remains constant, andso a higher coefficient of expansion points to an increase in freevolume associated with the system and therefore increased freedom forthe molecules to move. The increasing heat capacity corresponds to anincrease in heat dissipation through movement. The Tg of a given polymercan be dependent on the heating rate and can be influenced by thethermal history of the polymer. Furthermore, the chemical structure ofthe polymer heavily influences the glass transition by affectingmobility.

The transition from the brittle to ductile state is not sharp ordiscontinuous for an amorphous polymer. Rather, as the temperature ofthe polymer approaches Tg, the polymer becomes less brittle. Thesmoothness of the transition for an amorphous polymer can be illustratedby the temperature dependence of a number of properties. For example,FIG. 11 is a schematic plot of the specific volume of an amorphouspolymer vs. temperature. In the glassy and ductile states, the specificvolume vs. temperature is substantially linear, with each state having adifferent slope. As FIG. 11 shows, as the temperature is increased inthe linear glassy region, at a temperature T_(NL) the slope becomesnonlinear through a continuous transition region ΔT_(t). As shown, theTg is determined by the intersection of the slope of the specific volumevs. temperature in the brittle and ductile states.

The specific volume data depicted in FIG. 11 for an amorphous polymercan be obtained in a dilatometer. In this apparatus, a sample is placedin a glass bulb and a confining liquid, usually mercury, is introducedinto the bulb so that the liquid surrounds the sample and extendspartway up a narrow bore glass capillary tube. A capillary tube is usedso that relatively small changes in polymer volume caused by changingthe temperature produce easily measured changes in the height of themercury in the capillary.

The determination of Tg for amorphous materials, such as polymers bydilatometric methods (as well as by other methods) are found to be ratedependent. For example, a higher cooling rate can result in a highervalue of Tg. Tg can also be determined using differential scanningcalorimetry. Other methods can also be used that use measurements ofquantities such as density, dielectric constant, and elastic modulus.

The increase in the specific volume is associated with increasedmolecular motion which can result in less brittle behavior. Therefore, apolymer deformed in this region is less susceptible to mechanicalfailure. Additionally, a deformed polymer is less likely to recoiltowards the undeformed state. This reduction in the polymers desire toreturn to an undeformed state is due to the large scale segmental chainmotion that is possible for the polymer above Tg. This increased chainmotion will result in very low residual stress in the deformed state,which reduces the desire for the polymer to return to an undeformedstate. For polymers deformed below Tg, there will be some retainedstress in the polymer, which will result in a desire for the polymer toreturn to the undeformed state. Therefore, it can be desirable for astent to be at a temperature above the nonlinear transition, T_(NL),during crimping. In some embodiments, a polymeric stent may be at atemperature between T_(NL) and Tg, or more narrowly in a rangeT_(NL)+0.90*ΔT_(t), T_(NL)+0.70*ΔT_(t), T_(NL)+0.50*ΔT_(t),T_(NL)+0.30*ΔT_(t), or T_(NL)+0.10*ΔT_(t), during crimping.

In further embodiments, the stent may be crimped at a selectedtemperature Ts, which is less than Tg. Other embodiments includecrimping in a temperature interval Ts−ΔT_(s)<T<Ts+ΔT_(s). For example,Ts may be Tg−25° C. and ΔT_(s) may be 5° C., 15° C., or 20° C.

Furthermore, some embodiments may include a stent at a temperature at orabove Tg during crimping. However, crimping at such a temperature shouldbe done in a manner that reduces or prevents loss of inducedorientation. The higher the molecular weight of the polymer and/orfaster the crimping process the lower the loss of orientation.

A device for crimping the polymeric stent can resemble any crimpingdevice as is known in the art. Additionally, the device can beespecially modified so that it can heat the stent during crimping. Insome embodiments, the device can apply pressure and heat simultaneously.In these or other embodiments, after crimping, the crimping device canhold the stent at an elevated temperature, which may be selected suchthat it is greater than, equal to, or less than the selected crimpingtemperature or may be selected to specifically exclude temperaturesgreater than, equal to, or less than the selected crimping temperature.In some embodiments, the device crimps the polymeric stent while thestent is heated by other means.

The stent can be heated for up to one hour, 30 seconds to one hour, orfor 30 seconds. In some embodiments, the stent is heated long enoughthat the material becomes ductile enough to adequately lower thebrittleness of the stent. Adequate means having a value for theparameter in question such that one of ordinary skill in the art wouldexpect the invention to function in the particular application. Forexample, “adequately lower the brittleness of the stent” means that thebrittleness of the stent is reduced enough to warrant the extra heatingstep and the extra cost and complication of the heating step, as viewedby one of ordinary skill in the art.

For the purposes of the present invention, the following terms anddefinitions apply:

“Stress” refers to force per unit area, as in the force acting through asmall area within a plane. Stress can be divided into components, normaland parallel to the plane, called normal stress and shear stress,respectively. Tensile stress, for example, is a normal component ofstress applied that leads to expansion (increase in length). Inaddition, compressive stress is a normal component of stress applied tomaterials resulting in their compaction (decrease in length). Stress mayresult in deformation of a material, which refers to a change in length.“Expansion” or “compression” may be defined as the increase or decreasein length of a sample of material when the sample is subjected tostress.

“Strain” refers to the amount of expansion or compression that occurs ina material at a given stress or load. Strain may be expressed as afraction or percentage of the original length, i.e., the change inlength divided by the original length. Strain, therefore, is positivefor expansion and negative for compression.

“Strength” refers to the maximum stress along an axis which a materialwill withstand prior to fracture. The ultimate strength is calculatedfrom the maximum load applied during the test divided by the originalcross-sectional area.

“Modulus” may be defined as the ratio of a component of stress or forceper unit area applied to a material divided by the strain along an axisof applied force that results from the applied force. For example, amaterial has both a tensile and a compressive modulus.

The tensile stress on a material may be increased until it reaches a“tensile strength” which refers to the maximum tensile stress which amaterial will withstand prior to fracture. The ultimate tensile strengthis calculated from the maximum load applied during a test divided by theoriginal cross-sectional area. Similarly, “compressive strength” is thecapacity of a material to withstand axially directed pushing forces.When the limit of compressive strength is reached, a material iscrushed.

“Toughness” is the amount of energy absorbed prior to fracture, orequivalently, the amount of work required to fracture a material. Onemeasure of toughness is the area under a stress-strain curve from zerostrain to the strain at fracture. The units of toughness in this caseare in energy per unit volume of material. See, e.g., L. H. Van Vlack,“Elements of Materials Science and Engineering,” pp. 270-271,Addison-Wesley (Reading, Pa., 1989).

The underlying structure or substrate of an implantable medical device,such as a stent can be completely or at least in part made from abiodegradable polymer or combination of biodegradable polymers, abiostable polymer or combination of biostable polymers, or a combinationof biodegradable and biostable polymers. Additionally, a polymer-basedcoating for a surface of a device can be a biodegradable polymer orcombination of biodegradable polymers, a biostable polymer orcombination of biostable polymers, or a combination of biodegradable andbiostable polymers.

It is understood that after the process of degradation, erosion,absorption, and/or resorption has been completed, no part of the stentwill remain or in the case of coating applications on a biostablescaffolding, no polymer will remain on the device. In some embodiments,very negligible traces or residue may be left behind. For stents madefrom a biodegradable polymer, the stent is intended to remain in thebody for a duration of time until its intended function of, for example,maintaining vascular patency and/or drug delivery is accomplished.

Representative examples of polymers that may be used to fabricate orcoat an implantable medical device include, but are not limited to,poly(N-acetylglucosamine) (Chitin), Chitosan, poly(hydroxyvalerate),poly(lactide-co-glycolide), poly(hydroxybutyrate),poly(hydroxybutyrate-co-valerate), polyorthoester, polyanhydride,poly(glycolic acid), poly(glycolide), poly(L-lactic acid),poly(L-lactide), poly(D,L-lactic acid), poly(D,L-lactide),poly(caprolactone), poly(trimethylene carbonate), polyester amide,poly(glycolic acid-co-trimethylene carbonate), co-poly(ether-esters)(e.g. PEO/PLA), polyphosphazenes, biomolecules (such as fibrin,fibrinogen, cellulose, starch, collagen and hyaluronic acid),polyurethanes, silicones, polyesters, polyolefins, polyisobutylene andethylene-alphaolefin copolymers, acrylic polymers and copolymers otherthan polyacrylates, vinyl halide polymers and copolymers (such aspolyvinyl chloride), polyvinyl ethers (such as polyvinyl methyl ether),polyvinylidene halides (such as polyvinylidene chloride),polyacrylonitrile, polyvinyl ketones, polyvinyl aromatics (such aspolystyrene), polyvinyl esters (such as polyvinyl acetate),acrylonitrile-styrene copolymers, ABS resins, polyamides (such as Nylon66 and polycaprolactam), polycarbonates, polyoxymethylenes, polyimides,polyethers, polyurethanes, rayon, rayon-triacetate, cellulose, celluloseacetate, cellulose butyrate, cellulose acetate butyrate, cellophane,cellulose nitrate, cellulose propionate, cellulose ethers, andcarboxymethyl cellulose. Another type of polymer based on poly(lacticacid) that can be used includes graft copolymers, and block copolymers,such as AB block-copolymers (“diblock-copolymers”) or ABAblock-copolymers (“triblock-copolymers”), or mixtures thereof.

Additional representative examples of polymers that may be especiallywell suited for use in fabricating or coating an implantable medicaldevice include ethylene vinyl alcohol copolymer (commonly known by thegeneric name EVOH or by the trade name EVAL), poly(butyl methacrylate),poly(vinylidene fluoride-co-hexafluororpropene) (e.g., SOLEF 21508,available from Solvay Solexis PVDF, Thorofare, N.J.), polyvinylidenefluoride (otherwise known as KYNAR, available from ATOFINA Chemicals,Philadelphia, Pa.), ethylene-vinyl acetate copolymers, and polyethyleneglycol.

EXAMPLES

Some embodiments of the present invention are illustrated by thefollowing prophetic Examples. The Examples are being given by way ofillustration only and not by way of limitation. The parameters and dataare not be construed to unduly limit the scope of the embodiments of theinvention.

Embodiments of a stent can be fabricated from poly(L-lactide) (PLLA).

Tube Manufacturing

A polymeric tube can be fabricated using a 1 in single screw extruder.The temperature of the polymer melt within the extruder can be between370° F. and 470° F., 390° F. and 450° F., or more narrowly, between 310°F. and 430° F. The residence time in extruder can be less than 15minutes, less than 10 minutes, or more narrowly, less than 5 minutes.The tube exiting the die can be cooled or quenched in a water bath. Thelength of the quench or cooling distance can be less than 2 inches, lessthan 1 inch, or more narrowly less than ¾ inch. The pull rate of thetube from the die can be less than 25 ft/min, 20 ft/min, or morenarrowly less than 16 ft/min. The barrel pressure inside the extrudercan be between 1800 and 2200 psi. The axial draw down ratio of the tubeexiting the die can be approximately 3:1.

The pre-extrusion and post extrusion number average molecular weight(Mn), weight average molecular weight (Mw), and polydispersity (PD) canbe as follows:

-   -   Pre-extrusion—Mn can be between 255 K and 275 K, Mw can be        between 450 510 K and 530 K, and PD can be between 1.9 and 2    -   Post-extrusion—Mn can be between 170 K and 190 K, Mw can be        between 370 K and 390 K, and PD can be between 2 and 2.1.

The post-extrusion % crystallinity can be between 10% and 15%.

The post-extrusion strain to failure (measured at 98.6° F. (37° C.) anda stretch rate of 0.5 inch/min) can be between 290% and 310%.

Radial Expansion

The extruded tubing can be expanded from 0.018 in inside diameter(ID)/0.056 in outside diameter (OD) to (0.065 in to 0.080 in ID)/(0.077in to 0.092 in OD), with 30-60% longitudinal stretch of the tube. Thetubing can be expanded by blow molding, as described above, into a glassmold. The degree of crystallinity after expansion can be between 35% and55%. The temperature of tube during radial expansion can be between 160°F. and 180° F.

The pre-blow molding and post blow molding Mn, Mw, and PD can be asfollows:

-   -   Pre-blow molding—Mn can be between 175 K and 185 K, Mw can be        between 375 K and 385 K, and PD can be between 2 and 2.1.    -   Post-blow molding—Mn can be between 160 K and 170 K, Mw can be        between 335 K and 345 K, and PD can be between 1.9 and 2.

Laser Machining

Laser machining can be performed with a laser having a 120 femto secondpulse. The wavelength of the laser can be 800 nm. The pre-laser cuttingand post laser cutting Mn, Mw, and PD can be as follows:

Pre-laser cutting—Mn can be between 160 K and 170 K, Mw can be between335 K and 345 K, and PD can between 1.9 and 2.

Post laser cutting—Mn can be between 115 K and 125 K, Mw can be between275 K and 285 K, and PD can between 2.25 and 2.35.

Stent Pattern

A stent pattern as depicted in FIG. 9 can be cut with the laser. Stentstruts can have a rectangular or square cross-section. For example, thestruts can measure 0.006×0.0065 in (150×150 micron).

Crimping

As indicated above, the stent can be crimped from the cut diameter to adesired diameter onto a support element, such as a balloon. A slidingwedge style crimper can be used. The crimp cycle may be between about 25and 30 seconds. The stent can be heated to a temperature between 28° C.and 32° C. during crimping. The stent can be crimped from a 0.084 in ODto a 0.053 in OD.

Deployment

The crimped stent can be deployed with an outward radial pressure in theballoon of 7 atm to 0.118 (3.0 mm) ID or a pressure of 21 atm to 0.138in (3.5 mm) ID.

While particular embodiments of the present invention have been shownand described, it will be obvious to those skilled in the art thatchanges and modifications can be made without departing from thisinvention in its broader aspects. Therefore, the appended claims are toencompass within their scope all such changes and modifications as fallwithin the true spirit and scope of this invention.

1-15. (canceled)
 16. A method of fabricating a stent comprising: forminga polymer tube made of a biodegradable polymer, wherein thebiodegradable polymer of the formed tube is amorphous or substantiallyamorphous; processing the tube to increase the crystallinity of thebiodegradable polymer from amorphous or substantially amorphous tobetween 35% and 55%; cutting a pattern in the processed tube to form acylindrically-shaped scaffold, wherein the scaffold includes a patternof interconnecting struts, wherein the pattern comprises a plurality ofcylindrical rings of struts and longitudinal linking struts connectingthe rings, wherein the rings include bending elements including strutsthat flex inward to allow crimping of the scaffold and flex outward toallow expansion of the scaffold, wherein the scaffold is radiallyexpandable in a blood vessel of a body and has adequate radial strengthto hold open the blood vessel.
 17. The method of claim 16, wherein thebiodegradable polymer is a copolymer.
 18. The method of claim 16,wherein the biodegradable polymer is a mixture of polymers.
 18. Themethod of claim 16, wherein the processing comprises increasing atemperature of the tube above a glass transition temperature (Tg) of thebiodegradable polymer.
 19. The method of claim 16, wherein theprocessing comprises radial expansion of the tube to the target diameterat a temperature above the Tg of the biodegradable polymer.
 20. Themethod of claim 19, wherein a radial draw ratio of the radial expansionof the tube is between 2 and
 6. 21. The method of claim 16, furthercomprising crimping the stent on to a balloon at a diameter less thanthe target diameter so that the stent can be delivered into the bloodvessel.